1. Field of the Invention
This invention relates to motion correction for a radiation imaging device, such as a scintillation camera, to overcome image blurring due to movement of the object under study.
2. Description of the Prior Art
Radiation imaging devices are widely used as diagnostic tools for analyzing the distribution of a radiation-emitting substance in an object under study, such as in the medical diagnosis of a human body organ.
A typical radiation imaging device of a type to which the present invention is applicable is an Anger-type scintillation camera, the basic principles of which are described in U.S. Pat. No. 3,011,057. In such a device, quanta of radiation emitted by the object under study impinge upon a transducer in the form of a scintillation crystal to cause scintillation events. The events are detected by photodetectors arranged behind the transducer and analyzed to develop event X, Y positional coordinate signals representative of the relative spatial origins of the various quanta. The energies of the events are also analyzed to develop unblanking signals, so that only those events within a specified energy range of interest will be recorded for imaging, such as by display on a cathode ray tube oscilloscope. The mechanism for developing positional signals and unblanking signals for radiation imaging devices is well known and is described, for example, in U.S. Pat. No. 3,984,689.
The usefulness of a radiation imaging device as a diagnostic tool depends on the ability of the device to faithfully reproduce the distribution in the object under study of the radiation-emitting substance. One factor limiting this ability is the motion of the object, viz. the patient or body organ, during the diagnostic procedure. For imaging onto a CRT display, such motion produces undesirable blurring with consequential image deterioration. Use of a CRT display is common in medical diagnoses, such as liver scans, heart studies and the like.
Attempts to reduce the effects of patient movement during radiation imaging procedures have included the use of physical restraints to keep the patient still. Such devices are of limited value, cause patient discomfort and limit the types of tests that can be performed.
A more acceptable approach to motion correction for radiation imaging devices is the use of compensating circuitry which normalizes the positional signals developed from the detected radiation events in order to overcome the effects of object motion. Previously proposed motion correction methods of this type utilize the positional signals developed for display of events within the desired energy range, to calculate the mean or centroid position of the object under study. The positional signals are then corrected for object motion in accordance with the monitored displacement of the calculated object centroid.
A motion correction scheme using a digital computer on liver scans to correct for blurring due to patient breathing is described in Oppenheim, "A Method Using a Digital Computer for Reducing Respiratory Artifact on Liver Scans Made with a Camera," J. Nucl. Med., Vol. 12, pages 625-628 (1971). A computer memory was used to store event data in matrix form. The median row of the data matrix was determined for each 0.6 second data collection interval. Each matrix was then shifted so that all image frames had their median rows at the same location. The various shifted data frames were then summed, producing a resultant image for which the blurring effects of the shifting object median position were minimized. This procedure was relatively expensive, complex and required increased processing time.
A more convenient and less expensive method of motion correction was proposed by Hoffer which used simple analog circuitry, as described in U.S. Pat. No. 3,780,290. A schematic of the Hoffer circuit is shown in the prior art FIG. 1 for the correction of the X positional coordinate signal. Duplicate circuitry is employed for correction of the Y positional coordinate signal.
With reference to prior art FIG. 1, incoming X positional coordinate signals are fed into an integrator comprising an operational amplifier A1, a capacitor C1 and two resistors R1, as shown. The magnitude of the signal determines the coordinate position of the corresponding radiation event along an X-axis direction. Signals with positive values, corresponding to events occurring on the +X side of the X-axis, serve to charge the integrating capacitor C1; signals with negative values, corresponding to events occurring on the -X side, serve to charge the integrating capacitor C1 with the opposite polarity. The voltage on the integrating capacitor C1 at any given time is the mean or centroid position X of the preceding distribution of events sampled by the integrator. A summer comprising an operational amplifier A2 and resistors R.sub.X, R.sub.X and R2, serves to normalize each succeeding positional signal by subtracting the calculated centroid position from the positional coordinate signal of each event to be displayed. For example, if the calculated centroid moves from a position X=0 to X=2, the value X=2 will be subtracted from the positional coordinate signals of the displayed events, resulting in a displayed image that is stationary on the viewing monitor, unblurred by the motion. This approach offers advantages over the digital approach in that it does not add to the image time or decrease the accumulated count rate.
Improved analog correction circuitry is described in McKeighen, "Improved Means of Correcting Blurring in Scintigraphic Images," Phys. Med. Biol., Vol. 2, No. 2, pages 353-362 (1979). Such circuitry is shown in prior art FIG. 2. As in the Hoffer circuitry (prior art FIG. 1), the displacement of the centroid position of the radiation distribution of detected events is used to normalize the positional signals of the successive events. The centroid integrator, comprising the amplifier A1, the capacitor C1 and the resistors R1, is modified to include a switch SW1 which is closed in response to a strobe signal. The strobe signal corresponds to the unblanking signal developed by event energy analyzer circuitry and causes the switch SW1 to close whenever an event has been detected which is within the specified energy range of events to be displayed. The centroid integrator of the circuitry in prior art FIG. 2 offers advantages over the centroid integrator circuitry of prior art FIG. 1. In Hoffer's circuit, each event charges the integrator until the next event comes in. Thus the pulse widths driving the integrator are variable from event to event and depend on the counting rate. Also, the circuit estimates the centroid based on the count rate. The circuitry of prior art FIG. 2, on the other hand, estimates the centroid on a fixed number of events, independent of count rate. Further, the feedback resistor R1 of the integrator of prior art FIG. 2 is only connected across the capacitor C1 during a pulse interval, which prevents the feedback resistor from bleeding the charge of the capacitor during intervals when pulses are not charging the integrator.
The circuitry of prior art FIG. 2 includes a second integrator for calculating and holding the initial position of the centroid. This "initial position hold" circuit comprises the amplifier A3, the integrating capacitor C3 and the resistors R3. A switch SW3 serves to gate positional signals to the integrating capacitor C3 for an initial constant time interval determined by front panel controls on the imaging device. The "initial position hold" circuit serves to sample the initial position of the object and hold it in its relative position on the display, regardless of subsequent motion. This offers an advantage over the Hoffer circuit of prior art FIG. 1 in which the centroid position is always taken to be X=0. The prior art circuitry of FIG. 2 also includes an input buffer comprising the amplifier A4 and the resistors R4.
The summer, comprising the amplifier A2 and resistors R2, R.sub.X, R.sub.X and R.sub.X0, adjusts the uncorrected positional signal by subtracting from it the displacement of the centroid from its initial position: X.sub.C =X-(X-X0).
The described circuits all have the disadvantage that the centroid displacement developed for correcting the displayed positional signals is based on the positional distribution of the displayed signals themselves. The use of centroids for motion correction which are based on emanations from the organ under study itself is undesirable because the centroids are influenced by factors which may be uncorrectable. An example of this is in the use of motion centroids to correct the patient motion in heart studies. Here the centroid is influenced by motion from respiration and intrinsic cardiac contraction, both of which adversely influence the correction. Twisting and turning of the heart itself (i.e. motion of the heart in a Z-axis direction) interferes with the calculation of the centroid of the distribution in the X-Y plane. The addition of respiration monitoring an electrocardiograph (ECG) gating does not wholly compensate for such distorting factors.